Ultrasound imaging uses sound waves with frequencies higher than those audible to humans (>20 000 Hz). These sound waves are pulsed into tissue using a probe. The sound waves echo off the tissue, with different tissues reflecting varying degrees of sound. These echoes are analyzed through signal processing and are further processed using clinical ultrasound reconstruction algorithms to reconstruct ultrasound images for presentation and interpretation by an operator. Many different types of images can be reconstructed using ultrasound imaging. One such type is a B-mode image which displays the acoustic impedance of a two-dimensional cross-section of tissue.
Ultrasound imaging reconstruction requires knowledge of the speed of sound in the imaged medium. Although the speed of sound in soft tissues ranges from 1450 m/s in fat to 1590 m/s in the liver, clinical ultrasound reconstruction algorithms assume a homogenous speed of sound of 1540 m/s which is a weighted average for a model human. Spatial variation of the speed of sound within a region of interest of the imaged medium and the discrepancies between assumed values and actual values result in phase and refractive errors or aberrations incurred for both the transmitted sound waves and the reflected echoes. These aberrations contribute to contrast and resolution loss in reconstructed images as well as introduce distortions, artifacts and errors in feature localization.
An exemplary ultrasound imaging system is shown in FIG. 1. As can be seen, the ultrasound imaging system comprises ultrasound transducer arrays 10 that emit sound waves into a region of interest (ROI), in this case tissue comprising non-fat regions 12 and 16 and fat region 14. As the sound waves travel from the non-fat region 12 to the fat region 14 to the non-fat region 16, the sound waves are refracted by an angle of θr. As a result, the position P of a structure within the region of interest ROI appears at shifted positions P′. The discrepancy between the actual position and the apparent position of the structure may be as high as several millimeters and significantly affects the resolution of the ultrasound image. If the thickness of fat region 14 is known, the shift in the position of the structure can be corrected. However, in obese patients, the thickness of the fat region can vary considerably resulting in blurred images.
A number of methods and techniques have been developed to correct for fat-induced aberrations in ultrasound imaging due to a spatially-varying speed of sound. In some common methods, algorithms are used that infer the average ultrasound phase distortion along a given direction by analyzing the arrival phase variations from known guide stars or beacons. These methods are however, computationally intensive, require a number of iterations and are not always accurate.
In other common methods, clinical information obtained using various imaging modalities, such as for example ultrasound imaging, X-ray computed tomography or magnetic resonance imaging, is relied upon to determine the boundaries and thickness of fat layers. Instead of using a homogenous speed of sound, these methods use two or more values to represent the speed of sound within each layer. These methods require access to expensive imaging modalities and are limited by their capability to delineate the boundaries of fat layers in obese patients.
Still other methods have been employed. For example, U.S. Pat. No. 6,705,994 to Vortman et al. discloses a method of imaging a site of interest in a body using an ultrasound probe comprising a plurality of ultrasound transducer elements. The method comprises obtaining an ultrasound image of a pass zone between the ultrasound probe and the site of interest. The image includes the site of interest and a plurality of tissue regions in the pass zone between the site of interest and the ultrasound probe. Boundaries of a selected tissue region in the pass zone are determined from the image. Focusing delay times are then computed for each ultrasound transducer element based in part on the speed of sound in the selected tissue region other than an average speed of sound in body tissue, and the boundaries of the selected tissue region. A speed of sound in the selected tissue region is used. In embodiments, refraction is considered. Ultrasound imaging of the site of interest is conducted employing the computed delay times. Fat and bone tissue regions are selected if present. Other tissue regions may be selected, as well. The boundaries of the tissue region or regions may be determined by segmentation. Tissue inhomogeneity is thereby compensated for, improving image contrast resolution.
As another example, U.S. Pat. No. 8,784,318 to Napolitano et al. discloses an ultrasound scanner equipped with an image data processing unit that can perform adaptive parameter optimization during image formation and processing. In an embodiment, an ultrasound system comprises a channel data memory to store channel data obtained by digitizing ultrasound image data produced by an image scan, an image data processor configured to process the stored channel data in the memory to reconstruct an ultrasound image for each of a plurality of trial values of at least one parameter to be optimized, and a parameter optimization unit configured to evaluate an image quality of the reconstructed ultrasound image for each trial value of the at least one parameter, and to determine the optimized value of the at least one parameter based on the evaluated image quality.
Although techniques for correcting for fat-induced aberrations in ultrasound imaging have been considered, improvements are desired. It is therefore an object at least to provide a novel method and system for correcting fat-induced aberrations in ultrasound imaging.